This invention relates to implantable medical devices such as defibrillators and automatic implantable defibrillators (AIDs), and their various components. More particularly, it relates to an implantable medical device including a single, flat battery configured to optimize an overall size and shape of the device.
Implantable medical devices for therapeutic stimulation of the heart are well known in the art. In U.S. Pat. No. 4,253,466 issued to Hartlaub et al., for example, a programmable demand pacemaker is disclosed. The demand pacemaker delivers electrical energy, typically ranging in magnitude between about 5 and about 25 micro Joules, to the heart to initiate the depolarization of cardiac tissue. This stimulating regime is used to treat the heart by providing pacemaker spike in the absence of naturally occurring spontaneous cardiac depolarizations.
Another form of implantable medical device for therapeutic stimulation of the heart is an automatic implantable defibrillator (AID), such as those described in U.S. Pat. No. Re. 27,757 to Mirowski et al. and U.S. Pat. No. 4,030,509 to Heilman et al. Those AID devices deliver energy (about 40 Joules) to the heart to interrupt ventricular fibrillation of the heart. In operation, an AID device detects the ventricular fibrillation and delivers a nonsynchronous high-voltage pulse to the heart through widely spaced electrodes located outside of the heart, thus mimicking transthoracic defibrillation. The technique of Heilman et al. requires both a limited thoracotomy to implant an electrode near the apex of the heart and a pervenous electrode system located in the superior vena cava of the heart. Another example of a prior art implantable cardioverter includes the pacemaker/cardioverter/defibrillator (PCD) disclosed in U.S. Pat. No. 4,375,817 to Engle et al. This device detects the onset of tachyarrhythmia and includes means to monitor or detect the progression of the tachyarrhythmia so that progressively greater energy levels may be applied to the heart to interrupt a ventricular tachycardia or fibrillation.
Another device is an external synchronized cardioverter, such as that described in xe2x80x9cClinical Application of Cardioversionxe2x80x9d in Cardiovascular Clinics, 1970, Vol. 2, pp. 239-260 by Douglas P. Zipes. This type of external device provides cardioversion shocks synchronized with ventricular depolarization to ensure that the cardioverting energy is not delivered during the vulnerable T-wave portion of the cardiac cycle.
Another example of a prior art implantable cardioverter includes the device disclosed in U.S. Pat. No. 4,384,585 to Douglas P. Zipes. This device includes circuitry to detect the intrinsic depolarizations of cardiac tissue and pulse generator circuitry to deliver moderate energy level stimuli (in the range of about 0.1 to about 10 Joules) to the heart synchronously with the detected cardiac activity.
The functional objective of such a stimulating regimen is to depolarize areas of the myocardium involved in the genesis and maintenance of re-entrant or automatic tachyarrhythmias at lower energy levels with greater safety than was possible with nonsynchronous cardioversion. Nonsynchronous cardioversion always incurs the risk of precipitating ventricular fibrillation and sudden death, Synchronous cardioversion delivers the shock at a time when the bulk of cardiac tissue is already depolarized and is in a refractory state. Other examples of automatic implantable synchronous cardioverters include those of Charms in U.S. Pat. No. 3,738,370.
It is expected that the increased safety deriving from use of lower energy levels and their attendant reduced trauma to the myocardium, as well as the smaller size of implantable medical devices, will expand indications for use beyond the existing patient base of automatic implantable defibrillators. Since many episodes of ventricular fibrillation are preceded by ventricular (and in some cases, supraventricular) tachycardias, prompt termination of the tachycardia may prevent ventricular fibrillation.
Consequently, current devices for the treatment of tachyarrhythmias include the possibility of programming staged therapies of antitachycardia pacing regimens, along with cardioversion energy and defibrillation energy shock regimens in order to terminate the arrhythmia with the most energy-efficient and least traumatic therapies, when possible. In addition, some current implantable tachycardia devices are capable of delivering single or dual chamber bradycardia pacing therapies, as of which are described, for example, in U.S. Pat. No. 4,800,833 to Winstrom, U.S. Pat. No. 4,830,006 to Haluska et al., and U.S. patent application Ser. No. 07/612,758 to Keimel for xe2x80x9cApparatus for Delivering Single and Multiple Cardioversion and Defibrillation Pulsesxe2x80x9d filed Nov. 14, 1990, and incorporated herein by reference in its entirety. Furthermore, and as described in the foregoing ""833 and ""006 patents and the ""758 application, considerable study has been undertaken to devise the most efficient electrode systems and shock therapies.
Initially, implantable cardioverters and defibrillators were envisioned as operating with a single pair of electrodes applied on or in the heart.
Examples of such systems are disclosed in the aforementioned ""757 and ""509 patents, wherein shocks are delivered between an electrode is placed in or on the right ventricle and a second electrode placed outside the right ventricle. Studies have indicated that two electrode defibrillation systems often require undesirably high-energy levels to effect defibrillation.
In an effort to reduce the amount of energy required to effect defibrillation, numerous suggestions have been made with regard to multiple electrode systems. Some of those suggestions are set forth in U.S. Pat. No. 4,291,699 to Geddes et al., U.S. Pat. No. 4,708,145 to Tacker et al., U.S. Pat. No. 4,727,877 to Kallock, and U.S. Pat. No. 4,932,407 issued to Williams where sequential pulse multiple electrode systems are described. Sequential pulse systems operate based on the assumption that sequential defibrillation pulses delivered between differing electrode pairs have an additive effect such that the overall energy requirements to achieve defibrillation are less than the energy levels required to accomplish defibrillation using a single pair of electrodes.
An alternative approach to multiple electrode sequential pulse defibrillation is disclosed in U.S. Pat. No. 4,641,656 to Smits and also in the above-cited ""407 patent. This defibrillation method may conveniently be referred to as a multiple electrode simultaneous pulse defibrillation method, and involves the simultaneous delivery of defibrillation pulses between two pairs of electrodes. For example, one electrode pair may include a right ventricular electrode and a coronary sinus electrode, and a second electrode pair may include a right ventricular electrode and a subcutaneous patch electrode, with the right ventricular electrode serving as a common electrode to both electrode pairs. An alternative multiple electrode, single path, biphasic pulse system is disclosed in U.S. Pat. No. 4,953,551 to Mehra et al., which employs right ventricular, superior vena cava and subcutaneous patch electrodes.
In the above-cited prior art simultaneous pulse multiple electrode systems, delivery of simultaneous defibrillation pulses is accomplished by simply coupling two electrodes together. For example, in the above-cited ""551 patent, the superior vena cava and subcutaneous patch electrodes are electrically coupled together and a pulse is delivered between those two electrodes and the right ventricular electrode. Similarly, in the above-cited ""407 patent, the subcutaneous patch and coronary sinus electrodes are electrically coupled together, and a pulse is delivered between these two electrodes and a right ventricular electrode. See also U.S. Pat. Nos. 5,411,539; 5,620,477; 5,6589,321; 5,545,189 and 5,578,062, where active can electrodes are discussed.
The aforementioned ""758 application discloses a pulse generator for use in conjunction with an implantable cardioverter/defibrillator which is capable of providing all three of the defibrillation pulse methods described above, with a minimum of control and switching circuitry. The output stage is provided with two separate output capacitors which are sequentially discharged during sequential pulse defibrillation and simultaneously discharged during single or simultaneous pulse defibrillation. The complexity of those stimulation therapy regimens require rapid and efficient charging of high voltage output capacitors from low voltage battery power sources incorporated within the implantable medical device.
Typically, the electrical energy required to power an implantable cardiac pacemaker is supplied by a low voltage, low current drain, long-lived power source such as a lithium iodine pacemaker battery of the type manufactured by Wilson Greatbatch, Ltd. or Medtronic, Inc. While the energy density of such power sources is typically relatively high, they are generally not capable of being rapidly and repeatedly discharged at high current drains in the manner required to directly cardiovert the heart with cardioversion energies in the range of 0.1 to 10 Joules. Moreover, the nominal voltage at which such batteries operate is generally too low for cardioversion applications. Higher energy density battery systems are known which can be more rapidly or more often discharged, such as lithium thionyl chloride power sources. Neither of the foregoing battery types, however, may have the capacity or the voltage required to provide an impulse of the required magnitude on a repeatable basis to the heart following the onset of tachyarrhythmia.
Generally speaking, it is necessary to employ a DC-DC converter to convert electrical energy from a low voltage, low current power supply to a high voltage energy level stored in a high-energy storage capacitor. A typical form of DC-DC converter is commonly referred to as a xe2x80x9cflybackxe2x80x9d converter which employs a transformer having a primary winding in series with the primary power supply and a secondary winding in series with the high-energy capacitor. An interrupting circuit or switch is placed in series with the primary coil and battery. Charging of the high-energy capacitor is accomplished by inducing a voltage in the primary winding of the transformer creating a magnetic field in the secondary winding. When the current in the primary winding is interrupted, the collapsing field develops a current in the secondary winding which is applied to the high-energy capacitor to charge it. The repeated interruption of the supply current charges the high-energy capacitor to a desired level over time.
In U.S. Pat. No. 4,548,209 to Wielders et al. and in the above-referenced ""883 patent, charging circuits are disclosed which employ flyback oscillator voltage converters which step up the power source voltage and apply charging current to output capacitors until the capacitor voltage reaches a programmed shock energy level.
In charging circuit 34 of FIG. 4 in the ""209 patent, two series-connected lithium thionyl chloride batteries 50 and 52 are connected to primary coil 54 of transformer 56 and to power FET transistor switch 60. Secondary coil 58 is connected through diode 62 to cardioversion energy storage capacitor 64. In this circuit, the flyback converter works generally as follows: When switch 60 is closed, current Ip passing through primary winding 54 increases linearly as a function of the formula Vp=Lpdl/dt. When FET 60 is opened, the flux in the core of transformer 56 cannot change instantaneously, and so complimentary current Is (which is proportional to the number of windings in primary and secondary coils 54 and 58, respectively) starts to flow in secondary winding 58 according to the formula Is=(NpNs)Ip. Simultaneously, voltage in the secondary winding is developed according to the function Vs=Lsdls/dt, thereby causing charging of cardioversion energy storage capacitor 64 to a programmed voltage.
The Power FET 60 is switched xe2x80x9conxe2x80x9d at a constant frequency of 32 KHz for a duration or duty cycle that varies as a function of the voltage of the output capacitor reflected back into the primary coil 54 circuit. The on-time of power FET 60 is governed by the time interval between the setting and resetting of flip-flop 70, which in turn is governed either by current Ip flowing through primary winding 54 or as a function of a time limit circuit containing further circuitry to vary the time limit with battery impedance (represented schematically by resistor 53). In both cases, the on time varies from a maximum to a minimum interval as the output circuit voltage increases to its maximum value.
The aforementioned ""883 and ""006 patents disclose a variable duty cycle flyback oscillator voltage converter, where the current in the primary coil circuit (in the case of the ""883 patent) or the voltage across a secondary coil (in the case of the ""006 patent) is monitored to control the duty cycle of the oscillator. In the ""883 circuit the xe2x80x9conxe2x80x9d time of the oscillator is constant and the xe2x80x9coffxe2x80x9d time varies as a function of the monitored current through the transformer.
In the ""006 patent, a secondary coil is added to power a high voltage regulator circuit that provides V+ to a timer circuit and components of the high voltage oscillator. This high voltage power source allows the oscillator circuit to operate independently of the battery source voltage (which may deplete over time). The inclusion of a further secondary winding on an already relatively bulky transformer is disadvantageous from size and efficiency standpoints.
Energy, volume, thickness and mass are critical features in the design of implantable cardiac defibrillators (ICDs). One of the components important to optimization of those features is the high voltage capacitor used to store the energy required for defibrillation. Such capacitors typically deliver energy in the range of about 25 to 40 Joules, while ICDs typically have a volume of about 40 to about 60 cc, a thickness of about 13 mm to about 16 mm and a mass of approximately 100 grams.
It is desirable to reduce the volume, thickness and mass of such capacitors and devices without reducing deliverable energy. Doing so is beneficial to patient comfort and minimizes complications due to erosion of tissue around the device. Reductions in size of the capacitors may also allow for the balanced addition of volume to the battery, thereby increasing longevity of the device, or balanced addition of new components, thereby adding functionality to the device. It is also desirable to provide such devices at low cost while retaining the highest level of performance.
Most ICDs employ commercial photoflash capacitors similar to those described by Troup in xe2x80x9cImplantable Cardioverters and Defibrillators,xe2x80x9d Current Problems in Cardiology, Volume XIV, Number 12, December 1989, Year Book Medical Publishers, Chicago, and U.S. Pat. No. 4,254,775 for xe2x80x9cImplantable Defibrillator and Package Thereforxe2x80x9d. The electrodes in such capacitors are typically spirally wound to form a coiled electrode assembly. Most commercial photoflash capacitors contain a core of separator paper intended to prevent brittle anode foils from fracturing during coiling. The anode, cathode and separator are typically wound around such a paper core. The core limits both the thinness and volume of the ICDs in which they are placed. The cylindrical shape of commercial photoflash capacitors also limits the volumetric packaging efficiency and thickness of an ICD made using same.
Recently developed flat aluminum electrolytic capacitors have overcome some disadvantages inherent in commercial cylindrical capacitors. For example, U.S. Pat. No. 5,131,388 to Pless et. al. discloses a relatively volumetrically efficient flat capacitor having a plurality of planar layers arranged in a stack. Each layer contains an anode layer, a cathode layer and means for separating the anode layers and cathode layers (such as paper). The anode layers and the cathode layers are electrically connected in parallel.
A segment of today""s ICD market employs flat capacitors to overcome some of the packaging and volume disadvantages associated with cylindrical photoflash capacitors. Examples of such flat capacitors are described in the ""388 patent to Pless et al. for xe2x80x9cImplantable Cardiac Defibrillator with Improved Capacitors,xe2x80x9d and the ""851 patent to Fayram for xe2x80x9cCapacitor for an Implantable Cardiac Defibrillatorsxe2x80x9d Additionally, flat capacitors are described in a paper entitled xe2x80x9cHigh Energy Density Capacitors for Implantable Defibrillatorsxe2x80x9d by P. Lunsmann and D. MacFarlane presented at the 16th Capacitor and Resistor Technology Symposium.
An additional design constraint relates to size and location of the energy source associated with an implantable medical device (IMD) for therapeutic stimulation of the heart. As explained in greater detail above, an IMD consists generally of a sealed housing maintaining a capacitor(s), an electronics module(s) and an energy source. The electronics module normally includes a circuit board maintaining a variety of electrical components designed, for example, to perform sensing and monitoring functions or routines, as well as to accumulate data related to IMD operation. The electronics module is electrically connected to the capacitor and the power source such that amongst other functions, the electronics module causes the power source to charge and recharge the capacitor. To satisfy power and safety requirements, the power source typically consists of two series-connected batteries. So as to optimize volumetric efficiency, the batteries are typically formed to assume a cube-like shape. For example, a well-accepted IMD configuration includes two, three-volt cube-like batteries connected in series.
Due to the preferred cube-like shape, the batteries must be positioned in a side-by-side fashion within the housing so as to minimize an overall thickness of the IMD. In other words, because the preferred design renders the batteries relatively thick, it is impractical to position one or both of the batteries over other components of the IMD. Such an arrangement would likely increase the overall IMD thickness beyond an acceptable level.
Numerous efforts have been made to improve upon the size, shape and performance characteristics of the various IMD components. For example, implementation of a flat capacitor configuration has greatly improved IMD performance as well as reducing and improving the size and shape of the IMD housing. Similarly, advancements in electrical component technology has greatly reduced size requirements associated with the electronics module, along with facilitating use of a lower voltage power source (e.g., three-volt versus six-volt). Along these same lines, enhancements in materials and construction techniques used for IMD batteries have resulted in the reduction of sizes and costs. Unfortunately, however, the generally accepted assembly approach of one or two batteries placed next to the capacitor and the electronics module has remained unchanged.
Manufacturers continue to improve upon IMD construction and size characteristics. To this end, currently available power source design and location is less than optimal. Therefore, a need exists for an IMD incorporating a power source unit having superior space-volumetric efficiencies to thereby advance the preferred objective for continuous IMD size reduction.
One aspect of the present invention provides an implantable medical device including a housing, a capacitor, an electronics module and a substantially flat battery. The capacitor is disposed within the housing. The electronics module is electrically connected to the capacitor and is disposed within the housing. More particularly, the electronics module generally defines opposing front and rear faces and is positioned such that the rear face is adjacent a wall of the housing. The battery is electrically connected to the electronics module. Further, the battery is positioned over the electronics module such that the battery is adjacent the front face of the electronics module. In one preferred embodiment, a single battery is provided, thereby reducing an overall housing volume compared to a dual battery configuration. Similarly, locating the substantially flat battery over the electronics module further minimizes required outer dimensions of the housing. In one preferred embodiment, the battery defines a form or shape factor corresponding with a form or shape factor of the electronics module.
Another aspect of the present invention relates to an implantable medical device including a housing, a capacitor, electronics module and a substantially flat battery. The housing includes opposing first and second shield walls. The capacitor is disposed within the housing. The electronics module is electrically connected to the capacitor and is disposed within the housing such that the electronics module is adjacent the first shield wall. The battery is electrically connected to the electronics module. Further, the battery is positioned within the housing between the electronics module and the second shield wall. With this configuration, an overall size of the housing is optimized. In one preferred embodiment, the battery and the electronics module combine to define a thickness corresponding with a thickness of the capacitor.
Yet another aspect of the present invention relates to a method of manufacturing an implantable medical device. The method includes providing a housing including opposing first and second shield walls. A capacitor is secured to the first shield wall. Similarly, an electronics module is secured to the first shield wall and electrically connected to the capacitor. A substantially flat battery is provided. The battery is placed over the electronics module and electrically connected thereto. Finally, the second shield wall is coupled to the first shield wall such that the battery is disposed within the housing between the electronics module and the second shield wall.
Yet another aspect of the present invention relates to a power source unit for use with an implantable medical device that includes a hermetically sealed enclosure maintaining an electronics module and a capacitor. The power source unit includes an electrochemical cell and a retainer. The electrochemical cell includes an anode, a cathode and an electrolyte contained within a case. The case has first and second opposing major surfaces. During use, the electrochemistry of the cell generates an internal swelling pressure within the case, for example in the form of a gas. The retainer is coupled to the first major surface and is configured to limit an outward deflection of the first major surface. With this configuration in mind, the retainer biases the case to deflect primarily along the second major surface in response to the internal swelling pressure. In one preferred embodiment, the retainer is a stiffener plate.